The use of artificial biomaterials is becoming increasingly widespread in several areas of medical treatment. For example, biomaterials are now used in the repair of damaged tissue (e.g. bone and skin), in prosthetic devices such as artificial hip and knee joints, heart valves, and blood vessels, and in drug delivery devices. However, one of the challenges that remains in this field of medicine is the provision of biomaterials with improved biocompatibility properties, which can be readily colonized by host cells.
The study of implant surface and biomaterial tissue interface reactions is essential for the continued improvement of implant performance. A review by Blitterwijk et al., (1991) discusses the importance of the reactions of cells at implant surfaces in determining the biocompatibility of the implant (i.e. how well the implant is accepted by the surrounding tissues and by the body as a whole). Current research is aimed at making ‘bioactive materials’ that will readily permit integration of the material into the host tissue.
There is considerable interest in poly(ε-caprolactone) (PCL) as a potential bioactive material. It has been widely used for the last thirty years for the production of resorbable sutures, biomedical implants, drug delivery systems and vaccine formulation. Currently, PCL is being exploited for bone graft substitution (Coombes and Meikle, 1994), because it could overcome the problems of limited supply of bone, risk of infection such as HIV, additional surgical operations and long bony union times associated with either autografts or allografts.
PCL is a semi-crystalline linear resorbable synthetic aliphatic polyester, —(—O—(CH2)5—CO—)—. When implanted in vivo, the polymer is readily degraded non-specifically by hydrolytic enzymes, esterases and carboxypeptidases. Pitt et al. (1981), showed that degradation of PCL in vivo and in vitro proceeds via hydrolytic chain scission of ester linkages until the segments are sufficiently small to diffuse through the polymer bulk. Once the polymer reaches the molecular weight of 5000 Daltons, significant weight loss is observed, which is dependent on particle size. Chain scission has also been shown to be associated with an increase in crystallinity, which partially determines the rate of degradation (see review by Smith et al., 1990).
The products generated from the degradation of PCL are either incorporated into the tricarboxylic acid (TCA) cycle and removed by the lungs in the form of carbon dioxide and water, or eliminated by direct renal secretion. Taylor et al. (1994) tested PCL in vitro for the acute toxicity of degradation products. They found that the pH of PCL in sterile distilled water and Tris buffer remained relatively constant over sixteen weeks, and that the samples degraded slightly more in Tris compared to in water. It has also been found that hydroxy radicals produced by inflammatory cells play a major role in the degradation of PCL in vitro (Ali et al., 1992, 1993) and in vivo (Ali et al., 1994). The bioresorbability and biocompatibility of PCL is reviewed by Vert et al., (1992).
Another favourable factor of PCL is that it can be blended with numerous other polymers, e.g. Poly(L-Lactide) (PLA), to produce co-polymers with optimised properties. PLA is one of the strongest polyesters and has a resorption time of greater than one year, probably in the range of 2-3 years. This would be advantageous, for example, in a 3-D tissue construct/scaffold, where the implant resorption rate needs to be adjusted according to the tissue repair rate.
Feng et al., (1983) synthesised a biodegradable block copolymer of poly(ε-caprolactone) with poly(DL-lactide). The copolymers possessed release properties similar to silicone rubber (one of the first non-degradable drug delivery systems) but their degradation rates were always faster than that of PCL or PLA homopolymers. They intended to combine the excellent permeability of PCL with the faster biodegradation rate of PLA. Pitt et al., (1979) did investigate PCL, PLA and their copolymers and demonstrated how variabilities in the permeability of the drug delivery system could be achieved using copolymers of PCL and PLA, because PCL is more permeable than PLA.
Jianzhong et al, 1995 used PCL and poly(ethylene glycol) (PEG) block copolymers as a drug release device. It was found that the increasing PEG content of the copolymer caused an increase in the hydrophilicity and a decrease in the crystallinity of the copolymer. Thus, the drug releas behaviour and the degradability of the copolymer can be controlled by adjusting the composition of the copolymer.
Chan and Pitt (1990) tested the degradability of PCL when fabricated by compression moulding, co-precipitation and solvent evaporation and found that the method of fabrication and the resulting morphology of the polymers plays a critical role in determining their relative rates of hydrolytic degradation. Compression moulding of PCL/polyglycolic acid-co-lactic acid blends, increased the rate of chain scission as compared to the other fabrication methods.
A problem with PCL is that it is a plastic at body temperature. Its mechanical properties make it ideal for drug delivery systems, but not for the internal fixation of bone. Lowry et al., (1997) made reinforced PCL with phosphate glass fibres in the form of intra-medullary pins for the internal fixation of bone. This study was performed in the rabbit model and histological evidence showed that the composite was well tolerated, with minimal inflammation around the pin. The review by Daniels et al. (1990) illustrates that the mechanical properties of polymers and composites can be improved by reinforcing the materials with alumina, alumina-boria-silica, calcium metaphosphate glass fibers and carbon.
As well as blending PCL with other synthetic polymers to control the degradation rate, improve the mechanical properties of the system and alter its permeability, PCL can also be blended with natural polymers, e.g. collagen, fibronectin, hyaluronic acid and glycosaminoglycans. These natural polymers are all part of the extracellular matrix (ECM) that cells produce and secrete. It is thought that by incorporating natural polymers with synthetic polymers, osteoconductance and biocompatibility properties could be combined with the physical and mechanical properties of the synthetic component, making bioartificial polymers a good bioactive biomaterial substitute.
Giusti et al., (1994) discuss the importance of blending collagen with polymeric materials for use as medical devices and show how blending also increases the mechanical and thermal properties as compared to the individual components. Cascone et al., (1994) demonstrated the use of collagen and hyaluronic acid based polymeric bioartificial polymers as a successful drug delivery system for the release of growth hormone.
Several reports have shown the importance of the ECM in cellular function Ruoslahti et al., 1985 and Ellis and Yannis, 1996). Cells initially attach to the biomaterial by physicochemical factors, e.g. charge, surface free energy or the water content of the biomaterial (Schamberger and Gardella, 1994), and then strongly adhere to ECM proteins, which have been deposited on the biomaterial surface. How the ECM is deposited, stabilised and configured on a particular biomaterial surface is still not known, however tissue transglutaminase (tTG) has been implicated in the stabilisation process. It is important to understand this process in order to control cellular responses to surfaces.
Transglutaminases (Enzyme Commission System of Classification 2.3.2.13) are a group of multifunctional enzymes that cross-link and stabilise proteins in tissues and body fluids (Aeschlimann and Paulsson, 1994 and Greenberg et al., 1991). In mammals, they are calcium dependent and catalyse the post-translational modification of proteins by forming inter and intra-molecular ε(γ-glutamyl) lysine cross-links. The bonds that form are stable, covalent and resistant to proteolysis, thereby increasing the resistance of tissues to chemical, enzymatic and physical disruption. In contrast to transglutaminases of mammalian origin, microbial transglutaminases are generally not Ca2+-dependent.
The number of proteins acting as glutaminyl substrates for transglutaminases is highly restricted since both the primary structure and conformation are critical. In contrast, the only requirement of the acyl-acceptor substrate is the presence of a suitable pi amine, e.g. the ε-amino group of peptide bound lysine residues and small primary amines. Different types of transglutaminase enzyme differ in their specificity for a given glutaminyl substrate. For example, the plasma transglutaminase blood coagulation factor XIIIa acts on a limited range of glutaminyl substrates compared to tissue (or type II) transglutaminase (tTG). Unlike Factor XIIIa, tTG also binds GTP and GDP, which is thought to be important in its regulation by Ca2+ (see Smethurst and Griffin, 1996). A further key difference between the types of transglutaminase is in their distribution.
Although tTG has been mainly described as a cytosolic enzyme and does not contain a typical hydrophobic leader sequence for secretion, it may be found both in the cytosol and membrane associated depending on the cell type. The biological function of tTG has yet to be determined. However, there is now increasing evidence to suggest that tTG can act at the cell surface, facilitating cell adhesion (Borge et al., 1996) and cell spreading (Jones et al., 1997) and the modification of the extracellular matrix (ECM) (Aeschlimann et al., 1995, Barsigian et al., 1991, Bowness et al., 1988 and Bendixen et al., 1993).
The ability of transglutaminase enzymes to cross-link proteins has been exploited in the development of biological glues for promoting adhesion between tissue surfaces. For example, biological adhesive compositions comprising a tissue transglutaminase are disclosed in WO 94/28949. These compositions also comprise a divalent metal ion co-factor, which plays a regulatory role in the functional activity of transglutaminase enzymes (see Casadio et al., 1999, Eur. J. Biochem. 262, 672-679).